Radiographic imaging device and computer readable medium

ABSTRACT

There is provided a radiographic imaging device including: a radiation detector including plural radiographic image acquisition pixels that are arranged in a matrix in an imaging region for capturing a radiographic image and that acquire image information representing the radiographic image by converting applied radiation into electric charges and storing the electric charges and plural radiation detection pixels that are arranged in the imaging region, that have mutually different characteristics, and that detect the applied radiation by converting the applied radiation into electric charges and storing the electric charges; and a detecting unit that uses the radiation detection pixels selectively according to the characteristics to detect a state of application of the radiation.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation application of InternationalApplication No. PCT/2011/073577, filed Oct. 13, 2011, which isincorporated herein by reference. Further, this application claimspriority from Japanese Patent Application No. 2010-240076, filed Oct.26, 2010, which is incorporated herein by reference.

BACKGROUND OF THE INVENTION

1. Technical Field

The present invention pertains to a radiographic imaging device and acomputer readable medium storing a program, and particularly relates toa radiographic imaging device that captures a radiographic imagerepresented by radiation that has passed through a part to be imaged andto a computer readable medium storing a program executed by theradiographic imaging device.

2. Background Art

In recent years, radiation detectors such as flat panel detectors (FPD),in which a radiation-sensitive layer is placed on a thin-film transistor(TFT) active matrix substrate and which can directly convert radiationinto digital data, have been put into practical use, and radiographicimaging devices that use these radiation detectors to captureradiographic images expressed by applied radiation have also been putinto practical use. There are different methods by which the radiationdetectors used in these radiographic imaging devices convert theradiation, and these include the indirect conversion method, in whichthe radiation is converted into light by a scintillator and thereafterthe light is converted into electric charges by a semiconductor layer ofphotodiodes or the like, and the direct conversion method, in which theradiation is converted into electric charges by a semiconductor layer ofamorphous selenium or the like, and in each method, there exist variousmaterials that can be used for the semiconductor layer.

Incidentally, in this type of radiographic imaging device, if the startof the application of the radiation, the stopping of the application ofthe radiation, and the applied dose of the radiation can be detected bythe radiographic imaging device itself, it becomes unnecessary toconnect the radiographic imaging device to the radiation source and theimaging controller that collectively controls the radiographic imagingdevice and the radiation source, so this is preferred for simplifyingthe system configuration and simplifying the control by the imagingcontroller.

As a technology relating to this kind of radiographic imaging devicethat can detect the state of application of the radiation, in JP-A No.H07-201490, there is disclosed an X-ray diagnostic system includes anX-ray-to-optical signal converting unit that converts X-rays intooptical signals, optical-to-electrical signal converting means thatcaptures, with plural pixels, the optical signals converted by theX-ray-to-optical signal converting means and converts the opticalsignals into electrical signals, and X-ray exposure dose controllingmeans that controls the X-ray exposure dose with the electrical signalvalues of some of the pixels of the optical-to-electrical convertingmeans.

Further, in JP-A No. 2005-223157, there is disclosed a radiographicimaging device having a radiographic image detector that detects aradiographic image of a subject, plural radiation dose detectors thatdetect the dose of radiation from the subject, and a controller thatdetermines the mode of utilizing the outputs of the plural radiationdose detectors on the basis of the placement of the radiographic imagingdevice.

Moreover, in JP-A No. 2004-170216, there is disclosed a radiationdetector having a radiation converter in which conversion elements thatconvert incident radiation into electrical signals are disposed on asubstrate, wherein the radiation converter has first pixels, in whichthe conversion elements are connected to signal lines via switchelements that transfer the electrical signals and which output signalsfor generating an image, and second pixels, in which the conversionelements are directly connected to the signal lines and which detect theapplication of the radiation.

SUMMARY OF INVENTION

However, in the technologies disclosed in JP-A No. H07-201490, JP-A No.2004-223157, and JP-A No. 2004-170216, although the state of applicationof the radiation can be detected by the devices themselves, depending onthe imaging conditions for capturing the radiographic image, it is notalways the case that the state of application of the radiation can besuitably detected.

That is, for example, in the case of capturing a radiographic imageusing only part of an imaging region resulting from the radiographicimaging device, such as when the part to be imaged is a leg, an arm, orthe like, ordinarily imaging is performed with the part to be imagedbeing positioned in the central portion of the imaging region. For thisreason, the levels of radiation doses that are detected differ greatlybetween the radiation dose obtained by radiation detection pixelsdisposed in the imaging region where the part to be imaged is notpositioned and the radiation dose obtained by radiation detection pixelsdisposed in the imaging region where the part to be imaged ispositioned, so in a case where the characteristics of each of theradiation detection pixels are fixed in common, there are cases wherethe radiation dose of one ends up becoming saturated and thesignal-to-noise ratio (S/N ratio) of the radiation dose of the otherends up being remarkably low.

Further, for example, in the case of capturing a moving radiographicimage, the radiation dose is reduced compared to the case of capturing astill radiographic image, but even in a case where the characteristicsof the radiation detection pixels are fixed in common between capturinga moving image and capturing a still image, there are cases where theradiation dose of one ends up becoming saturated and the signal-to-noiseratio (S/N ratio) of the radiation dose of the other ends up beingremarkably low.

The present invention provides a radiographic imaging device and acomputer readable medium storing a program that can detect a state ofapplication of radiation more accurately.

According to a first aspect of the present invention, there is provideda radiographic imaging device including: a radiation detector includingplural radiographic image acquisition pixels that are arranged in amatrix in an imaging region for capturing a radiographic image and thatacquire image information representing the radiographic image byconverting applied radiation into electric charges and storing theelectric charges and plural radiation detection pixels that are arrangedin the imaging region, that have mutually different characteristics, andthat detect the applied radiation by converting the applied radiationinto electric charges and storing the electric charges; and a detectingunit that uses the radiation detection pixels selectively according tothe mutually different characteristics to detect a state of applicationof the radiation.

According to the radiographic imaging device according to the firstaspect, the image information representing the radiographic image isacquired by the radiation detector as a result of the applied radiationbeing converted into electric charges and the electric charges beingstored by the plural radiographic image acquisition pixels arranged in amatrix in the imaging region for capturing the radiographic image.

Here, in the present invention, the plural radiation detection pixelsthat are arranged in the imaging region, have mutually differentcharacteristics, and detect the applied radiation by converting theapplied radiation into electric charges and storing the electric chargesare used selectively according to the mutually different characteristicsand the state of application of the radiation is detected by thedetecting unit.

In this way, according to the radiographic imaging device according tothe first aspect, the plural radiation detection pixels that havemutually different characteristics are arranged in the radiationdetector, and the radiation detection pixels are used selectivelyaccording to the mutually different characteristics to detect the stateof application of the radiation, therefore the state of application ofthe radiation can be detected more accurately compared to a case wherethe radiation detector does not have these pixels.

According to a second aspect of the present invention, in the firstaspect, the radiographic imaging device may further include an acquiringunit that acquires an imaging condition for capturing the radiographicimage, and the detecting unit may use the radiation detection pixelswhich have characteristics corresponding to the imaging conditionacquired by the acquiring unit to detect the state of application of theradiation. Because of this, the state of application of the radiationcan be detected more accurately compared to a case where the pixels areselected independently of the imaging condition.

Further, according to a third aspect of the present invention, in thefirst or second aspect, the radiation detection pixels may be arrangedin different positions in the imaging region. Because of this, the stateof application of the radiation can be detected more accurately as aresult of being able to select and use, in accordance with the size andso forth of the part to be imaged, the pixels used to detect the stateof application of the radiation.

Further, according to a fourth aspect of the present invention, in anyof the first to third aspects, the state of application of the radiationmay be at least one of a start of application of the radiation, an endof application of the radiation, or an applied dose of the radiation,and the detecting unit may select the radiation detection pixels used inthe detection of the state of application in accordance with the stateof application of the radiation to be detected. Because of this, thestate of application of the radiation can be detected more accurately asa result of being able to select and use, in accordance with the stateof application of the radiation to be detected, the pixels used todetect the state of application.

Further, according to a fifth aspect of the present invention, in any ofthe first to fourth aspects, the mutually different characteristics maybe different because the plural radiation detection pixels beingconnected to amplifiers that amplify, at mutually different gains,signals represented by the electric charges stored by the radiationdetection pixels. According to a sixth aspect of the present invention,in any of the first to fifth aspects, the mutually differentcharacteristics may be different because of the plural radiationdetection pixels being connected to low-pass filters that performlow-pass filtering at mutually different low-pass filtering frequencieswith respect to signals represented by the electric charges stored bythe radiation detection pixels. According to a seventh aspect of thepresent invention, in any of the first to sixth aspects, the mutuallydifferent characteristics may be different because of the pluralradiation detection pixels being connected to a synthesizing unit thatsynthesize mutually different numbers of signals represented by theelectric charges stored by the radiation detection pixels. Because ofthis, the characteristics of the pixels that have mutually differentcharacteristics can be realized easily.

Further, according to an eighth aspect of the present invention, in thesecond aspect, the imaging condition is at least one of a part to beimaged, a region in which the part to be imaged is placed when capturingthe radiographic image, whether the imaging is imaging to capture amoving image or a still image, or an applied dose of the radiation.Because of this, the state of application of the radiation can bedetected more accurately in accordance with the applied imagingcondition.

Moreover, according to a ninth aspect of the present invention, in anyof the first to eighth aspects, the radiation detector may be furtherincludes dedicated lines for reading out the stored electric chargesfrom the radiation detection pixels. Because of this, the radiographicimage can be captured at a higher speed as a result of being able todetect the state of application of the radiation independently of theaction of capturing the radiographic image.

According to a tenth aspect of the present invention, there is provideda computer readable medium storing a program executed by a radiographicimaging device including a radiation detector that includes pluralradiographic image acquisition pixels that are arranged in a matrix inan imaging region for capturing a radiographic image and that acquireimage information representing the radiographic image by convertingapplied radiation into electric charges and storing the electric chargesand plural radiation detection pixels that are arranged in the imagingregion, that have mutually different characteristics, and that detectthe applied radiation by converting the applied radiation into electriccharges and storing the electric charges, the program causing a computerto function as an acquiring unit that acquires an imaging condition forcapturing the radiographic image and a detecting unit that uses theradiation detection pixels having characteristics corresponding to theimaging condition acquired by the acquiring unit to detect a state ofapplication of the radiation.

Consequently, according to the invention according to the tenth aspect,a computer can be caused to act in the same way as the radiographicimaging device of the present invention, therefore, like theradiographic imaging device the state of application of the radiationcan be detected more accurately.

According to the present invention, the plural radiation detectionpixels that have mutually different characteristics are arranged in theradiation detector, and these pixels are used selectively according tothe characteristics to detect the state of application of the radiation,therefore the state of application of the radiation can be detected moreaccurately compared to a case where the radiation detector does not havethese pixels.

BRIEF DESCRIPTION OF DRAWINGS

FIG. 1 is a block diagram showing the configuration of a radiologyinformation system pertaining to an embodiment;

FIG. 2 is a side view showing an example of an arrangement of devices,in a radiographic imaging room, of a radiographic imaging systempertaining to the embodiment;

FIG. 3 is a cross-sectional schematic view showing the schematicconfiguration of three pixel sections of a radiation detector pertainingto the embodiment;

FIG. 4 is a cross-sectional side view schematically showing theconfiguration of a signal output portion of one pixel section of theradiation detector pertaining to the embodiment;

FIG. 5 is a plan view showing the configuration of the radiationdetector pertaining to the embodiment;

FIG. 6 is a plan view showing an arrangement of radiation detectionpixels pertaining to the embodiment;

FIG. 7 is a perspective view showing the configuration of an electroniccassette pertaining to the embodiment;

FIG. 8 is a cross-sectional side view showing the configuration of theelectronic cassette pertaining to the embodiment;

FIG. 9 is a block diagram showing the configurations of main portions ofan electrical system of the radiographic imaging system pertaining tothe embodiment;

FIG. 10 is a circuit diagram showing the configuration of a secondsignal processing unit pertaining to the embodiment;

FIG. 11 is a flowchart showing a flow of processing by a radiographicimaging processing program pertaining to the embodiment;

FIG. 12 is a schematic diagram showing an example of an initialinformation input screen pertaining to the embodiment;

FIG. 13 is a flowchart showing a flow of processing by a cassetteimaging processing program pertaining to the embodiment;

FIG. 14 is a flowchart showing a flow of processing by a first radiationdose acquisition processing routine program pertaining to theembodiment;

FIG. 15 is a flowchart showing a flow of processing by a secondradiation dose acquisition processing routine program pertaining to theembodiment;

FIG. 16 is a cross-sectional side view for describing an irradiationside sampling method and a penetration side sampling method for readinga radiographic image; and

FIG. 17 is a plan view showing another example of an arrangement of theradiation detection pixels pertaining to the embodiment.

DETAILED DESCRIPTION OF THE INVENTION

An embodiment of the present invention will be described in detail belowwith reference to the drawings. Here, an example configuration in a casein which the present invention is applied to a radiology informationsystem, which is a system that collectively manages information handledin a radiology department in a hospital, will be described.

First, the configuration of a radiology information system (hereinaftercalled “RIS”) 100 pertaining to the present embodiment will be describedwith reference to FIG. 1.

The RIS 100 is a system for managing information such as medical serviceappointments and diagnostic records in a radiology department andconfigures part of a hospital information system (hereinafter called“HIS”).

The RIS 100 has plural imaging request terminal devices (hereinaftercalled “terminal devices”) 140, an RIS server 150, and radiographicimaging systems (hereinafter called “imaging systems”) 104 installed inindividual radiographic imaging rooms (or operating rooms) in ahospital, and the RIS 100 is configured as a result of these beingconnected to an in-hospital network 102 including a wired or wirelesslocal area network (LAN) or the like. The RIS 100 configures part of theHIS disposed in the same hospital, and an HIS server (not shown in thedrawings) that manages the entire HIS is also connected to thein-hospital network 102.

The terminal devices 140 are for doctors or radiologic technologists toinput and browse diagnostic information and facility reservations, andradiographic imaging requests and imaging reservations are also made viathe terminal devices 140. Each terminal device 140 is configured toinclude a personal computer having a display device, and the terminaldevices 140 are capable of intercommunication with the RIS server 150via the in-hospital network 102.

The RIS server 150 receives the imaging requests from each of theterminal devices 140 and manages radiographic imaging schedules in theimaging systems 104, and the RIS server 150 is configured to include adatabase 150A.

The database 150A is configured to include: information relating topatients (subjects), such as attribute information (names, sexes, datesof birth, ages, blood types, body weights, patient identifications(IDs), etc.) of the patients, medical histories, consultation histories,radiographic images that have been captured in the past, etc.;information relating to later-described electronic cassettes 40 used inthe imaging systems 104, such as identification numbers (IDinformation), models, sizes, sensitivities, dates of first use, numbersof times used, etc.; and environment information representing theenvironments in which the electronic cassettes 40 are used to captureradiographic images, that is, the environments in which the electroniccassettes 40 are used (e.g., radiographic imaging rooms, operatingrooms, etc.).

The imaging systems 104 capture radiographic images as a result of beingoperated by the doctors or the radiologic technologists in response toan instruction from the RIS server 150. Each imaging system 104includes: a radiation generator 120 that applies a dose of radiation X(see also FIG. 7) according to exposure conditions from a radiationsource 121 (see also FIG. 2) to a subject; an electronic cassette 40that has a built-in radiation detector 20 (see also FIG. 7) that absorbsthe radiation X that has passed through the part of the subject to beimaged, generates electric charges, and creates image informationrepresenting a radiographic image on the basis of the generated electriccharge quantity; a cradle 130 that charges a battery built into theelectronic cassette 40; and a console 110 that controls the electroniccassette 40 and the radiation generator 120.

The console 110 acquires various types of information included in thedatabase 150A from the RIS server 150, stores the information in alater-described HDD 116 (see FIG. 9), uses the information as needed,and controls the electronic cassette 40 and the radiation generator 120.

In FIG. 2, there is shown an example of an arrangement of the devices,in a radiographic imaging room 180, of the imaging system 104 pertainingto the present embodiment.

As shown in FIG. 2, an upright-position stand 160 used when performingradiographic imaging in an upright position and a recumbent-positiontable 164 used when performing radiographic imaging in a recumbentposition are installed in the radiographic imaging room 180. The spacein front of the upright-position stand 160 serves as a subject imagingposition 170 when performing radiographic imaging in the uprightposition, and the space above the recumbent-position table 164 serves asa subject imaging position 172 when performing radiographic imaging inthe recumbent position.

A holder 162 that holds the electronic cassette 40 is disposed in theupright-position stand 160, and the electronic cassette 40 is held inthe holder 162 when capturing a radiographic image in the uprightposition. Likewise, a holder 166 that holds the electronic cassette 40is disposed in the recumbent-position table 164, and the electroniccassette 40 is held in the holder 166 when capturing a radiographicimage in the recumbent position.

Further, in order to enable both radiographic imaging in the uprightposition and radiographic imaging in the recumbent position withradiation from the single radiation source 121, a supporting and movingmechanism 124 that supports the radiation source 121 in such a way thatthe radiation source 121 is rotatable about a horizontal axis (thedirection of arrow a in FIG. 2), is movable in the vertical direction(the direction of arrow b in FIG. 2), and is movable in the horizontaldirection (the direction of arrow c in FIG. 2) is disposed in theradiographic imaging room 180. Here, the supporting and moving mechanism124 includes a drive source that rotates the radiation source 121 aboutthe horizontal axis, a drive source that moves the radiation source 121in the vertical direction, and a drive source that moves the radiationsource 121 in the horizontal direction (none of the drive sources areshown in the drawings).

An accommodating portion 130A capable of storing the electronic cassette40 is formed in the cradle 130.

When the electronic cassette 40 is not in use, the built-in battery ofthe electronic cassette 40 is charged by the cradle 130 with theelectronic cassette 40 stored in the accommodating portion 130A of thecradle 130, and when a radiographic image is to be captured, theelectronic cassette 40 is removed from the cradle 130 by a radiologictechnologist or the like and is held in the holder 162 of theupright-position stand 160 if the imaging posture is the uprightposition or is held in the holder 166 of the recumbent-position table164 if the imaging posture is the recumbent position.

Here, in the imaging system 104 pertaining to the present embodiment,various types of information are transmitted and received by wirelesscommunication between the radiation generator 120 and the console 110and between the electronic cassette 40 and the console 110.

The electronic cassette 40 is not used only in a state in which it isheld in the holder 162 of the upright-position stand 160 or the holder166 of the recumbent-position table 164, but because it is portable,when imaging an arm, a leg, or the like, it can also be used in a statein which it is not held in the holders.

Next, the configuration of the radiation detector 20 pertaining to thepresent embodiment will be described. FIG. 3 is a cross-sectionalschematic diagram schematically showing the configuration of three pixelsections of the radiation detector 20 pertaining to the presentembodiment.

As shown in FIG. 3, in the radiation detector 20 pertaining to thepresent embodiment, signal output portions 14, sensor portions 13, and ascintillator 8 are sequentially layered on an insulating substrate 1,and pixels are configured by the signal output portions 14 and thesensor portions 13. The pixels are plurally arrayed on the substrate 1and are configured in such a way that the signal output portion 14 andthe sensor portion 13 in each pixel lie on top of one another.

The scintillator 8 is formed via a transparent insulating film 7 on thesensor portions 13 and includes a phosphor film that converts radiationmade incident from above (the opposite side of the substrate 1 side) orbelow into light and luminesces. By disposing the scintillator 8, thescintillator 8 absorbs the radiation that has passed through the subjectand luminesces.

It is preferred that the wavelength range of the light emitted by thescintillator 8 be in the visible light range (a wavelength of 360 nm to830 nm), and it is more preferred that the wavelength range of the lightemitted by the scintillator 8 include the green wavelength range inorder to enable monochrome imaging by the radiation detector 20.

As the phosphor used for the scintillator 8, specifically a phosphorincluding cesium iodide (CsI) is preferred in the case of imaging usingX-rays as the radiation, and using CsI(Tl) (cesium iodide to whichthallium has been added) whose emission spectrum when X-rays are appliedis 420 nm to 700 nm is particularly preferred. The emission peakwavelength, in the visible light range, of CsI(Tl) is 565 nm. Thephosphor used for the scintillator is not limited to this, and GOS—andparticularly GOS:Tb(Gd202S:Tb) (terbium-activated gadoliniumoxysulfide)—or the like can also be used The emission peak wavelength,in the visible light range, of GOS:Tb is 550 nm.

The sensor portions 13 have an upper electrode 6, lower electrodes 2,and a photoelectric conversion film 4 that is placed between the upperand lower electrodes, and the photoelectric conversion film 4 isconfigured by an organic photoelectric conversion material that absorbsthe light emitted by the scintillator 8 and generates an electriccharge.

It is preferred that the upper electrode 6 be configured by a conductingmaterial that is transparent at least with respect to the emissionwavelength of the scintillator 8 because it is necessary to allow thelight produced by the scintillator 8 to be made incident on thephotoelectric conversion film 4; specifically, using a transparentconducting oxide (TCO) whose transmittance with respect to visible lightis high and whose resistance value is small is preferred. A metal thinfilm of Au or the like can also be used as the upper electrode 6, butits resistance value tends to increase when trying to obtain atransmittance of 90% or more, so TCO is more preferred. For example,ITO, IZO, AZO, FTO, SnO₂, TiO₂, ZnO₂, and so forth can be preferablyused, and ITO is most preferred from the standpoints of process ease,low resistance, and transparency. The upper electrode 6 may have asingle-layer configuration common to all the pixels or may be dividedper pixel.

The photoelectric conversion film 4 includes the organic photoelectricconversion material, absorbs the light emitted from the scintillator 8,and generates an electric charge corresponding to the absorbed light.The photoelectric conversion film 4 including the organic photoelectricconversion material in this way has a sharp absorption spectrum in thevisible range, virtually no electromagnetic waves other than the lightemitted by the scintillator 8 are absorbed by the photoelectricconversion film 4, and noise that is generated as a result of radiationsuch as X-rays being absorbed by the photoelectric conversion film 4 canbe effectively suppressed.

It is preferred that the absorption peak wavelength of the organicphotoelectric conversion material configuring the photoelectricconversion film 4 be as close as possible to the emission peakwavelength of the scintillator 8 so that the organic photoelectricconversion material most efficiently absorbs the light emitted by thescintillator 8. It is ideal that the absorption peak wavelength of theorganic photoelectric conversion material and the emission peakwavelength of the scintillator 8 coincide, but as long as the differencebetween both is small, the organic photoelectric conversion material cansufficiently absorb the light emitted from the scintillator 8.Specifically, it is preferred that the difference between the absorptionpeak wavelength of the organic photoelectric conversion material and theemission peak wavelength of the scintillator 8 with respect to radiationbe within 10 nm, and it is more preferred that the difference be within5 nm.

Examples of organic photoelectric conversion materials capable ofsatisfying this condition include quinacridone organic compounds andphthalocyanine organic compounds. For example, the absorption peakwavelength, in the visible range, of quinacridone is 560 nm, so ifquinacridone is used as the organic photoelectric conversion materialand CsI(Tl) is used as the material of the scintillator 8, it becomespossible to make the difference between the peak wavelengths within 5 nmand the amount of electric charge generated in the photoelectricconversion film 4 can be substantially maximized. Also in the case ofusing GOS:Tb as the material of the scintillator 8, it is possible tomake the difference between the peak wavelengths with quinacridone asthe organic photoelectric conversion material about 10 nm, and theamount of electric charge generated by the photoelectric conversion film4 can be substantially maximized.

Next, the photoelectric conversion film 4 applicable to the radiationdetector 20 pertaining to the present embodiment will be specificallydescribed.

The electromagnetic wave absorption/photoelectric conversion site in theradiation detector 20 pertaining to the present embodiment can beconfigured by the pair of electrodes 2 and 6 and an organic layer thatincludes the organic photoelectric conversion film 4 sandwiched betweenthe electrodes 2 and 6. More specifically, this organic layer can beformed by stacking or mixing together a site that absorbselectromagnetic waves, a photoelectric conversion site, anelectron-transporting site, a hole-transporting site, anelectron-blocking site, a hole-blocking site, a crystallizationpreventing site, electrodes, and an interlayer contact improving site.

It is preferred that the organic layer includes an organic p-typecompound or an organic n-type compound.

Organic p-type semiconductors (compounds) are donor organicsemiconductors (compounds) represented mainly by hole-transportingorganic compounds, are organic compounds having the property that theyeasily donate electrons, and more specifically are organic compoundswhose ionization potential is the smaller of the two when two organicmaterials are brought into contact with each other and used.Consequently, any organic compound is usable as the donor organiccompound provided that it is an electron-donating organic compound.

Organic n-type semiconductors (compounds) are accepter organicsemiconductors (compounds) represented mainly by electron-transportingorganic compounds, are organic compounds having the property that theyeasily accept electrons, and more specifically are organic compoundswhose electron affinity is the greater of the two when two organiccompounds are brought into contact with each other and used.Consequently, any organic compound is usable as the accepter organiccompound provided that it is an electron-accepting organic compound.

Materials applicable as the organic p-type semiconductor and the organicn-type semiconductor, and the configuration of the photoelectricconversion film 4, are described in detail in JP-A No. 2009-32854, sodescription will be omitted. The photoelectric conversion film 4 mayalso be formed so as to further include fullerenes or carbon nanotubes.

It is preferred that the film thickness of the photoelectric conversionfilm 4 be as large as possible in terms of absorbing the light from thescintillator 8, but if the film thickness of the photoelectricconversion film 4 becomes thicker than a certain extent, the strength ofthe electric field generated in the photoelectric conversion film 4 bythe bias voltage applied from both ends of the photoelectric conversionfilm 4 drops and the electric charges become unable to be collected, sothe film thickness is preferably from 30 nm to 300 nm, more preferablyfrom 50 nm to 250 nm, and particularly preferably from 80 nm to 200 nm.

In the radiation detector 20 shown in FIG. 3, the photoelectricconversion film 4 has a single-layer configuration common to all thepixels, but the photoelectric conversion film 4 may also be divided perpixel.

The lower electrodes 2 are thin films divided per pixel. The lowerelectrodes 2 can be configured by a transparent or opaque conductingmaterial, and aluminum, silver, and so forth can be suitably used.

The thickness of the lower electrodes 2 can be 30 nm to 300 nm, forexample.

In the sensor portions 13, one from the electric charge (holes,electrons) generated in the photoelectric conversion film 4 can be movedto the upper electrode 6 and the other can be moved to the lowerelectrodes 2 as a result of a predetermined bias voltage being appliedbetween the upper electrode 6 and the lower electrodes 2. In theradiation detector 20 of the present embodiment, a wire is connected tothe upper electrode 6, and the bias voltage is applied to the upperelectrode 6 via this wire. Further, the polarity of the bias voltage isdecided in such a way that the electrons generated in the photoelectricconversion film 4 move to the upper electrode 6 and the holes move tothe lower electrodes 2, but this polarity may also be the opposite.

It suffices for the sensor portions 13 configuring each of the pixels toinclude at least the lower electrodes 2, the photoelectric conversionfilm 4, and the upper electrode 6, but in order to suppress an increasein dark current, disposing at least either of an electron-blocking film3 and a hole-blocking film 5 is preferred, and disposing both is morepreferred.

The electron-blocking film 3 can be disposed between the lowerelectrodes 2 and the photoelectric conversion film 4 and can suppresselectrons from being injected from the lower electrodes 2 into thephotoelectric conversion film 4 and dark current from ending upincreasing when the bias voltage has been applied between the lowerelectrodes 2 and the upper electrode 6.

Electron-donating organic materials can be used for theelectron-blocking film 3.

It suffices for the material that is actually used for theelectron-blocking film 3 to be selected in accordance with, for example,the material of the adjacent electrodes and the material of the adjacentphotoelectric conversion film 4, and a material whose electron affinity(Ea) is greater by 1.3 eV or more than the work function (Wf) of thematerial of the adjacent electrodes and has an ionization potential (Ip)equal to or smaller than the ionization potential of the material of theadjacent photoelectric conversion film 4 is preferred. Materialsapplicable as the electron-donating organic material are described indetail in JP-A No. 2009-32854, so description will be omitted.

In order to allow the electron-blocking film 3 to reliably exhibit adark current suppressing effect and to prevent a drop in thephotoelectric conversion efficiency of the sensor portions 13, thethickness of the electron-blocking film 3 is preferably from 10 nm to200 nm, more preferably from 30 nm to 150 nm, and particularlypreferably from 50 nm to 100 nm.

The hole-blocking film 5 can be disposed between the photoelectricconversion film 4 and the upper electrode 6 and can suppress holes frombeing injected from the upper electrode 6 into the photoelectricconversion film 4 and dark current from ending up increasing when thebias voltage has been applied between the lower electrodes 2 and theupper electrode 6.

Electron-accepting organic materials can be used for the hole-blockingfilm 5.

In order to allow the hole-blocking film 5 to reliably exhibit a darkcurrent suppressing effect and to prevent a drop in the photoelectricconversion efficiency of the sensor portions 13, the thickness of thehole-blocking film 5 is preferably from 10 nm to 200 nm, more preferablyfrom 30 nm to 150 nm, and particularly preferably from 50 nm to 100 nm.

It suffices for the material that is actually used for the hole-blockingfilm 5 to be selected in accordance with, for example, the material ofthe adjacent electrode and the material of the adjacent photoelectricconversion film 4, and a material whose ionization potential (Ip) isgreater by 1.3 eV or more than the work function (Wf) of the material ofthe adjacent electrode and has an electron affinity (Ea) equal to orgreater than the electron affinity of the material of the adjacentphotoelectric conversion film 4 is preferred. Materials applicable asthe electron-accepting organic material are described in detail in JP-ANo. 2009-32854, so description will be omitted.

In a case where the bias voltage is set in such a way that, from theelectric charge generated in the photoelectric conversion film 4, theholes move to the upper electrode 6 and the electrons move to the lowerelectrode 2, the positions of the electron-blocking film 3 and thehole-blocking film 5 may be reversed. Further, the electron-blockingfilm 3 and the hole-blocking film 5 do not both have to be disposed; acertain degree of a dark current suppressing effect can be obtained aslong as either is disposed.

The signal output portions 14 are formed on the surface of the substrate1 below the lower electrodes 2 of each of the pixels. In FIG. 4, theconfiguration of the signal output portions 14 is schematically shown.

As shown in FIG. 4, in each of the signal output portions 14 pertainingto the present embodiment, a capacitor 9 that stores the electric chargethat has moved to the lower electrode 2 and a field-effect thin-filmtransistor (TFT) (hereinafter sometimes simply called a thin-filmtransistor) 10 that converts the electric charge stored in the capacitor9 into an electrical signal and outputs the electrical signal are formedin correspondence to the lower electrode 2. The region in which thecapacitor 9 and the thin-film transistor 10 are formed has a sectionthat coincides with the lower electrode 2 as seen in a plan view, and bygiving the signal output portion 14 this configuration, the signaloutput portion 14 and the sensor portion 13 lie on top of one another inthe thickness direction. In order to minimize the plane area of theradiation detector 20 (the pixels), it is preferred that the region inwhich the capacitor 9 and the thin-film transistor 10 are formed becompletely covered by the lower electrode 2.

The capacitor 9 is electrically connected to the corresponding lowerelectrode 2 via a wire of a conductive material that is formedpenetrating an insulating film 11 disposed between the substrate 1 andthe lower electrode 2. Because of this, the electric charge trapped inthe lower electrode 2 can be moved to the capacitor 9.

In the thin-film transistor 10, a gate electrode 15, a gate insulatingfilm 16, and an active layer (channel layer) 17 are layered, andmoreover, a source electrode 18 and a drain electrode 19 are formed apredetermined spacing apart from each other on the active layer 17.

The active layer 17 can, for example, be formed by amorphous silicon, anamorphous oxide, an organic semiconductor material, carbon nanotubes, orthe like. The material configuring the active layer 17 is not limited tothese.

As the amorphous oxide configuring the active layer 17, oxides includingat least one of In, Ga, and Zn (e.g., In—O amorphous oxides) arepreferred, oxides including at least two of In, Ga, and Zn (e.g.,In—Zn—O amorphous oxides, In—Ga—O amorphous oxides, or Ga—Zn—O amorphousoxides) are more preferred, and oxides including In, Ga, and Zn areparticularly preferred. As an In—Ga—Zn—O amorphous oxide, an amorphousoxide whose composition in a crystalline state is expressed byInGaO₃(ZnO)_(m) (where m is a natural number less than 6) is preferred,and particularly InGaZnO₄ is more preferred.

Examples of organic semiconductor materials capable of configuring theactive layer 17 include phthalocyanine compounds, pentacene, and vanadylphthalocyanine, but the organic semiconductor materials are not limitedto these. Configurations of phthalocyanine compounds are described indetail in JP-A No. 2009-212389, so description will be omitted.

By forming the active layer 17 of the thin-film transistor 10 from anamorphous oxide, an organic semiconductor material, or carbon nanotubes,the active layer 17 does not absorb radiation such as X-rays, or if itdoes absorb any radiation the amount is an extremely minute amount, sothe generation of noise in the signal output portion 14 can beeffectively suppressed.

Further, in a case where the active layer 17 is formed with carbonnanotubes, the switching speed of the thin-film transistor 10 can beincreased, and the thin-film transistor 10 can be formed having a lowdegree of absorption of light in the visible light range. In the case offorming the active layer 17 with carbon nanotubes, the performance ofthe thin-film transistor 10 drops significantly simply by mixing aninfinitesimal amount of a metal impurity into the active layer 17, so itis necessary to separate, extract, and form extremely high-purity carbonnanotubes by centrifugal separation or the like.

Here, the amorphous oxide, organic semiconductor material, or carbonnanotubes configuring the active layer 17 of the thin-film transistor 10and the organic photoelectric conversion material configuring thephotoelectric conversion film 4 are all capable of being formed intofilms at a low temperature. Consequently, the substrate 1 is not limitedto a substrate with high heat resistance, such as a semiconductorsubstrate, a quartz substrate, or a glass substrate, and a plastic orother flexible substrate, aramids, or bionanofibers can also be used.Specifically, polyester, such as polyethylene terephthalate,polybutylene phthalate, and polyethylene naphthalate, polystyrene,polycarbonate, polyethersulphone, polyarylate, polyimide, polycyclicolefin, norbornene resin, and poly(chloro-trifluoro-ethylene) or otherflexible substrates can be used. By employing a flexible substrate madeof plastic, the substrate can be made lightweight, which is advantageousfor portability, for example.

Further, an insulating layer for ensuring insulation, a gas barrierlayer for preventing the transmission of moisture and oxygen, anundercoat layer for improving flatness or adhesion to the electrodes orthe like, and other layers may also be disposed on the substrate 1.

High-temperature processes of 200 degrees or higher can be applied toaramids, so a transparent electrode material can be hardened at a hightemperature and given a low resistance, and aramids can also accommodateautomatic mounting of driver ICs including solder reflow processes.Further, aramids have a thermal expansion coefficient that is close tothat of indium tin oxide (ITO) or a glass substrate, so they have littlewarping after manufacture and do not break easily. Further, aramids canalso form a thinner substrate compared to a glass substrate or the like.An ultrathin glass substrate and an aramid may also be layered to form asubstrate.

Further, bionanoflbers are composites of cellulose microfibril bundles(bacterial cellulose) that a bacterium (Acetobacter xylinum) producesand a transparent resin. Cellulose microfibril bundles have a width of50 nm, which is a size that is 1/10 with respect to visible lightwavelengths, and have high strength, high elasticity, and low thermalexpansion. By impregnating and hardening a transparent resin such as anacrylic resin or an epoxy resin in bacterial cellulose, bionanoflbersexhibiting a light transmittance of about 90% at a wavelength of 500 nmwhile including fibers at 60 to 70% can be obtained. Bionanofibers havea low thermal expansion coefficient (3 to 7 ppm) comparable to siliconcrystal, a strength comparable to steel (460 MPa), high elasticity (30GPa), and are flexible, so they can form a thinner substrate 1 comparedto a glass substrate or the like.

In the present embodiment, a TFT substrate 30 is formed by sequentiallyforming the signal output portions 14, the sensor portions 13, and thetransparent insulating film 7 on the substrate 1, and the radiationdetector 20 is formed by adhering the scintillator 8 onto the TFTsubstrate 30 using, for example, an adhesive resin whose lightabsorbance is low.

As shown in FIG. 5, on the TFT substrate 30, pixels 32 configured toinclude the sensor portions 13, the capacitors 9, and the thin-filmtransistors 10 are plurally disposed two-dimensionally in one direction(a row direction in FIG. 5) and an intersecting direction (a columndirection in FIG. 5) with respect to the one direction.

Further, plural gate lines 34 that are disposed extending in the onedirection (the row direction) and are for switching on and off thethin-film transistors 10 and plural data lines 36 that are disposedextending in the intersecting direction (the column direction) and arefor reading out the electric charges via the thin-film transistors 10 inan on-state are disposed in the radiation detector 20.

The radiation detector 20 is formed in a tabular, quadrilateral shapehaving four sides on its outer edges in a plan view; more specifically,the radiation detector 20 is formed in a rectangular shape.

Here, in the radiation detector 20 pertaining to the present embodiment,some of the pixels 32 are used for detecting the state of application ofthe radiation, and the remaining pixels 32 capture radiographic images.Hereinafter, the pixels 32 for detecting the state of application of theradiation will be called radiation detection pixels 32A, and theremaining pixels 32 will be called radiographic image acquisition pixels32B.

The radiation detector 20 pertaining to the present embodiment cannotobtain pixel information of radiographic images in the positions wherethe radiation detection pixels 32A are arranged because the radiationdetector 20 captures radiographic images with the radiographic imageacquisition pixels 32B excluding the radiation detection pixels 32A ofthe pixels 32. For this reason, in the radiation detector 20 pertainingto the present embodiment, the radiation detection pixels 32A arearranged in such a way as to be dispersed, and the console 110 executesdefective pixel correction processing created by interpolating pixelinformation of radiographic images in the positions where the radiationdetection pixels 32A are arranged using pixel information that has beenobtained by the radiographic image acquisition pixels 32B positionedaround those radiation detection pixels 32A.

Here, the imaging system 104 pertaining to the present embodimentperforms imaging with the part to be imaged having been positioned atleast in the central portion of the imaging region in the case ofperforming imaging using the entire imaging region resulting from theradiation detector 20, such as a case where the part to be imaged is anabdomen or the like, and in the case of performing imaging using onlypart of the imaging region resulting from the radiation detector 20,such as a case where the part to be imaged is a leg, an arm, a hand, orthe like.

In the radiation detector 20 pertaining to the present embodiment, asschematically shown in FIG. 6 for example, the radiation detectionpixels 32A are arranged in regions (hereinafter called “central portiondetection regions”) 20A and 20B in the neighborhood of the centralportion of the imaging region of the radiation detector 20 and regions(hereinafter called “peripheral edge portion detection regions”) 20C to20F in the neighborhoods of the four corners of the peripheral edgeportion of the imaging region.

Additionally, in order to detect the state of application of theradiation, the electronic cassette 40 pertaining to the presentembodiment is disposed with a radiation dose acquisition function foracquiring information (hereinafter called “radiation dose information”)indicating the applied dose of the radiation X from the radiation source121.

For this reason, in the radiation detector 20 pertaining to the presentembodiment, as shown in FIG. 5, direct read-out lines 38, to whichconnecting portions between the capacitors 9 and the thin-filmtransistors 10 in the radiation detection pixels 32A are connected andwhich are for directly reading out the electric charges stored in thosecapacitors 9, are disposed extending in the one direction (the rowdirection). In the radiation detector 20 pertaining to the presentembodiment, one direct read-out line 38 is allocated with respect toplural radiation detection pixels 32A arranged side by side in the onedirection, and the connecting portions between the capacitors 9 and thethin-film transistors 10 in those plural radiation detection pixels 32Aare connected to a common (single) direct read-out line 38. The directread-out lines correspond to dedicated lines in the claims.

Next, the configuration of the electronic cassette 40 pertaining to thepresent embodiment will be described. In FIG. 7, there is shown aperspective view showing the configuration of the electronic cassette 40pertaining to the present embodiment.

As shown in FIG. 7, the electronic cassette 40 pertaining to the presentembodiment includes a casing 41 including a material that allowsradiation to pass through, and the electronic cassette 40 is given awaterproof and airtight structure. When the electronic cassette 40 isused in an operating room or the like, there is the concern that bloodor other contaminants may adhere to the electronic cassette 40.Therefore, by giving the electronic cassette 40 a waterproof andairtight structure and sterilizing the electronic cassette 40 as needed,one electronic cassette 40 can be used repeatedly.

A space A that accommodates various parts is formed inside the casing41, and the radiation detector 20 that detects the radiation X that haspassed through the subject and a lead plate 43 that absorbs backscatterrays of the radiation X are sequentially disposed inside the space Afrom an irradiated surface side of the casing 41 to which the radiationX is applied.

Here, in the electronic cassette 40 pertaining to the presentembodiment, the region of one surface of the tabular shape of the casing41 corresponding to the position where the radiation detector 20 isdisposed is a quadrilateral imaging region 41A capable of detecting theradiation. The surface of the casing 41 having the imaging region 41A isa top panel 41B of the electronic cassette 40, and in the electroniccassette 40 pertaining to the present embodiment, the radiation detector20 is placed in such a way that the TFT substrate 30 is on the top panel41B side, and the radiation detector 20 is adhered to the inside surfaceof the top panel 41B (the surface of the top panel 41B on the oppositeside of the surface on which the radiation is made incident) in thecasing 41.

As shown in FIG. 7, a case 42 that accommodates a later-describedcassette controller 58 and power source 70 (see FIG. 9 for both) isplaced on one end side of the inside of the casing 41 in a position thatdoes not coincide with the radiation detector 20 (outside the range ofthe imaging region 41A). The cassette controller 58 corresponds to anacquiring unit and a detecting unit in the claims.

The casing 41 is configured by carbon fiber, aluminum, magnesium,bionanofibers (cellulose microfibrils), or a composite material, forexample, in order to make the entire electronic cassette 40 lightweight.

As the composite material, for example, a material including reinforcedfiber resin is used, and carbon, cellulose, or the like is included inthe reinforced fiber resin. Specifically, as the composite material,carbon fiber reinforced plastic (CFRP), a composite material with astructure where a foam material is sandwiched by CFRP, or a compositematerial where the surface of a foam material is coated with CFRP isused. In the present embodiment, a composite material with a structurewhere a foam material is sandwiched by CFRP is used. Because of this,the strength (rigidity) of the casing 41 can be raised compared to acase where the casing 41 is configured by a carbon alone.

As shown in FIG. 8, inside the casing 41, supports 44 are disposed onthe inner surface of a back surface portion 41C opposing the top panel41B, and the radiation detector 20 and the lead plate 43 are placed sideby side in this order in the application direction of the radiation Xbetween the supports 44 and the top panel 41B. The supports 44 areconfigured by a foam material, for example, from the standpoint ofreducing weight and the standpoint of absorbing dimensional deviations,and the supports 44 support the lead plate 43.

As shown in FIG. 8, adhesive members 80 that detachably adhere the TFTsubstrate 30 of the radiation detector 20 are disposed on the innersurface of the top panel 41B. Double-sided tape, for example, is used asthe adhesive members 80. In this case, the double-sided tape is formedin such a way that the adhesive force of one adhesive surface isstronger than the adhesive force of the other adhesive surface.

Specifically, the surface whose adhesive force is weak (the weakadhesive surface) is set to have a 180° peel strength equal to or lessthan 1.0 N/cm. Additionally, the surface whose adhesive force is strong(the strong adhesive surface) contacts the top panel 41B, and the weakadhesive surface contacts the TFT substrate 30. Because of this, thethickness of the electronic cassette 40 can be made thinner compared toa case where the radiation detector 20 is fixed to the top panel 41B by,for example, fixing members such as screws. Further, even if the toppanel 41B deforms due to an impact or a load, the radiation detector 20follows the deformation of the top panel 41B whose rigidity is high, soonly large curvature (a gentle curve) arises and the potential for theradiation detector 20 to break due to localized low curvature becomeslower. Moreover, the radiation detector 20 contributes to improving therigidity of the top panel 41B.

In this way, in the electronic cassette 40 pertaining to the presentembodiment, the radiation detector 20 is adhered to the inside of thetop panel 41B of the casing 41, so the casing 41 is made separable intotwo between the top panel 41B side and the back surface portion 41Cside, and when the radiation detector 20 is adhered to the top panel 41Bor when the radiation detector 20 is detached from the top panel 41B,the casing 41 is separated into two between the top panel 41B side andthe back surface portion 41C side.

In the present embodiment, the adhesion of the radiation detector 20 tothe top panel 41B does not have to be performed in a clean room or thelike. The reason is because, in a case where foreign materials such asmetal fragments that absorb radiation have become mixed in between theradiation detector 20 and the top panel 41B, the foreign materials canbe removed by detaching the radiation detector 20 from the top panel41B.

Next, the configurations of relevant portions of an electrical system ofthe imaging system 104 pertaining to the present embodiment will bedescribed with reference to FIG. 9.

As shown in FIG. 9, in the radiation detector 20 built into theelectronic cassette 40, a gate line driver 52 is placed on one side oftwo sides adjacent to each other, and a first signal processing unit 54is placed on the other side. The individual gate lines 34 of the TFTsubstrate 30 are connected to the gate line driver 52, and theindividual data lines 36 of the TFT substrate 30 are connected to thefirst signal processing unit 54.

Further, an image memory 56, a cassette controller 58, and a wirelesscommunication unit 60 are disposed inside the casing 41.

The thin-film transistors 10 of the TFT substrate 30 are sequentiallyswitched on in row units by signals supplied via the gate lines 34 fromthe gate line driver 52. The electric charges that have been read out bythe thin-film transistors 10 switched to an on-state are transmittedthrough the data lines 36 as electrical signals and are input to thefirst signal processing unit 54. Because of this, the electric chargesare sequentially read out in row units, and a two-dimensionalradiographic image becomes acquirable.

Although it is not shown in the drawings, the first signal processingunit 54 includes amplifier circuits that amplify the input electricalsignals and sample-and-hold circuits for each of the individual datalines 36, and the electrical signals that have been transmitted throughthe individual data lines 36 are amplified by the amplifier circuits andare thereafter held in the sample-and-hold circuits. Further, amultiplexer and an analog-to-digital (A/D) converter are sequentiallyconnected to the output sides of the sample-and-hold circuits, and theelectrical signals held in the individual sample-and-hold circuits aresequentially (serially) input to the multiplexer and are converted intodigital image data by the A/D converter.

The image memory 56 is connected to the first signal processing unit 54,and the image data that have been output from the A/D converter of thefirst signal processing unit 54 are sequentially stored in the imagememory 56. The image memory 56 has a storage capacity capable of storinga predetermined number of frames of image data, and each timeradiographic imaging is performed, the image data obtained by theimaging are sequentially stored in the image memory 56.

The image memory 56 is also connected to the cassette controller 58. Thecassette controller 58 is configured to include a microcomputer,includes a central processing unit (CPU) 58A, a memory 58B including aread-only memory (ROM) and a random access memory (RAM), and anonvolatile storage unit 58C including a flash memory or the like, andcontrols the actions of the entire electronic cassette 40.

Moreover, the wireless communication unit 60 is connected to thecassette controller 58. The wireless communication unit 60 is adapted toa wireless local area network (LAN) standard represented by IEEE(Institute of Electrical and Electronics Engineers) 802.11a/b/g or thelike and controls the transmission of various types of informationbetween the electronic cassette 40 and external devices by wirelesscommunication. Via the wireless communication unit 60, the cassettecontroller 58 is made capable of wireless communication with an externaldevice such as the console 110 that performs control relating toradiographic imaging and is made capable of transmitting and receivingvarious types of information to and from the console 110 and the like.

Further, a power source 70 is disposed in the electronic cassette 40,and the various circuits and elements described above (the gate linedriver 54, the first signal processing unit 54, the image memory 56, thewireless communication unit 60, the microcomputer functioning as thecassette controller 58, etc.) operate on power supplied from the powersource 70. The power source 70 has a built-in battery (a rechargeablesecondary battery) so as to not impair the portability of the electroniccassette 40, and the power source 70 supplies power to the variouscircuits and elements from the charged battery. In FIG. 9, illustrationof wires connecting the various circuits and elements to the powersource 70 is omitted.

In the radiation detector 20 pertaining to the present embodiment, asecond signal processing unit 55 is placed on the opposite side of thegate line driver 52 across the TFT substrate 30 in order to realize theradiation dose acquisition function, and the individual direct read-outlines 38 of the TFT substrate 30 are connected to the second signalprocessing unit 55.

Here, the configuration of the second signal processing unit 55pertaining to the present embodiment will be described. In FIG. 10,there is shown a circuit diagram showing the configuration of the secondsignal processing unit 55 pertaining to the present embodiment.

As shown in FIG. 10, the second signal processing unit 55 pertaining tothe present embodiment includes variable gain pre-amps (charge amps) 92,low-pass filters (LPFs) 96, and sample-and-hold circuits 97 incorrespondence to each of the direct read-out lines 38 connected to theradiation detection pixels 32A disposed in the central portion detectionregion 20A. The variable gain pre-amps (charge amps) 92 correspond toamplifiers in the claims, and the low-pass filters (LPFs) 96 correspondto low-pass filters in the claims.

Each of the variable gain pre-amps 92 is configured to include an op-amp92A whose positive input side is grounded, a capacitor 92B that isconnected in parallel between the negative input side and the outputside of the op-amp 92A, and a reset switch 92F. The reset switches 92Fare switched by the cassette controller 58. Further, each of the LPFs 96is configured to include a resistor 96A and a capacitor 96C.

The second signal processing unit 55 pertaining to the presentembodiment also includes variable gain pre-amps 92, LPFs 96′ whoselow-pass frequency differs from that of the LPFs 96, and sample-and-holdcircuits 97 in correspondence to each of the direct read-out lines 38connected to the radiation detection pixels 32A disposed in the centralportion detection region 20B. In the present embodiment, low-passfilters whose low-pass frequency is lower than that of the LPFs 96 areapplied as the LPFs 96′. The LPFs (low-pass filters) 96′ correspond tolow-pass filters in the claims.

Further, the second signal processing unit 55 pertaining to the presentembodiment also includes variable gain pre-amps 92′ whose gain differsfrom that of the variable gain pre-amps 92, LPFs 96, and sample-and-holdcircuits 97 in correspondence to each of the direct read-out lines 38connected to the radiation detection pixels 32A disposed in theperipheral edge portion detection regions 20D and 20E. In the presentembodiment, variable gain pre-amps whose gain is higher than that of thevariable gain pre-amps 92 are applied as the variable gain pre-amps 92′.The variable gain pre-amps 92′ correspond to amplifiers in the claims.

Moreover, the second signal processing unit 55 pertaining to the presentembodiment also includes variable gain pre-amps 92, binning components94, LPFs 96, and sample-and-hold circuits 97 in correspondence to eachof the direct read-out lines 38 connected to the radiation detectionpixels 32A disposed in the peripheral edge portion detection regions 20Cand 20F. The binning components 94 synthesize, into one electricalsignal, the electrical signals that have been output from apredetermined number (in the present embodiment, two) of the directread-out lines 38. The binning components 94 correspond to asynthesizing unit in the claims.

Additionally, the second signal processing unit 55 pertaining to thepresent embodiment includes one multiplexer 98 and one A/D converter 99each. The sample timings of the sample-and-hold circuits 97 and outputsselected by switches 98A disposed in the multiplexer 98 are alsoswitched by the cassette controller 58.

The direct read-out lines 38 connected to the radiation detection pixels32A disposed in the central portion detection regions 20A and 20B andthe peripheral edge portion detection regions 20D and 20E areindividually connected to the input ends of the multiplexer 98 via theorder of the corresponding variable gain pre-amps 92 or variable gainpre-amps 92′, LPFs 96 or LPFs 96′, and sample-and-hold circuits 97.

Further, the direct read-out lines 38 connected to the radiationdetection pixels 32A disposed in the peripheral edge portion detectionregions 20C and 20F are individually connected to the input ends of themultiplexer 98 via the order of the corresponding variable gain pre-amps92, binning components 94, LPFs 96, and sample-and-hold circuits 97.Additionally, the output ends of the multiplexer 98 are connected to theinput end of the A/D converter 99, and the output end of the A/Dconverter 99 is connected to the cassette controller 58.

When causing the radiation dose acquisition function to work, thecassette controller 58 first discharges (resets) the electric chargesstored in the capacitors 92B by switching the reset switches 92F of thevariable gain pre-amps 92 and 92′ to an on-state for a predeterminedperiod of time.

The electric charges stored in the capacitors 9 of the radiationdetection pixels 32A as a result of the radiation X being applied aretransmitted as electrical signals through the connected direct read-outlines 38. The electrical signals transmitted through the direct read-outlines 38 are amplified at predetermined gains by the correspondingvariable gain pre-amps 92 and 92′, thereafter the electrical signalscorresponding to the radiation detection pixels 32A disposed in theperipheral edge portion detection regions 20C and 20F are synthesized bythe binning components 94, and filtering is performed at predeterminedlow-pass frequencies by the LPFs 96 and 96′.

After performing the discharge (reset), the cassette controller 58causes the sample-and-hold circuits 97 to hold the signal levels of thefiltered electrical signals by driving the sample-and-hold circuits 97for a predetermined period of time.

Then, the signal levels held in the sample-and-hold circuits 97 aresequentially selected by the multiplexer 98 in accordance with thecontrol by the cassette controller 58 and are converted from analog todigital by the A/D converter 99; thereafter, the digital signalsobtained thereby are output to the cassette controller 58. The digitalsignals output from the A/D converter 99 represent the dose of radiationthat was applied in the predetermined period of time with respect to theradiation detection pixels 32A and correspond to the radiation doseinformation.

Additionally, the cassette controller 58 sequentially stores, in apredetermined region of the RAM in the memory 58B, the radiation doseinformation that has been input from the A/D converter 99.

As shown in FIG. 9, the console 110 is configured as a server computerand includes a display 111, which displays operation menus, capturedradiographic images, and so forth, and an operation panel 112, which isconfigured to include plural keys and to which various types ofinformation and operation instructions are input.

Further, the console 110 pertaining to the present embodiment includes aCPU 113 that controls the actions of the entire device, a ROM 114 inwhich various programs including a control program are storedbeforehand, a RAM 115 that temporarily stores various types of data, ahard disk drive (HDD) 116 that stores and holds various types of data, adisplay driver 117 that controls the display of various types ofinformation on the display 111, and an operation input detector 118 thatdetects states of operation with respect to the operation panel 112.Further, the console 110 includes a wireless communication unit 119 thattransmits and receives various types of information such aslater-described exposure conditions to and from the radiation generator120 by wireless communication and also transmits and receives varioustypes of information such as image data to and from the electroniccassette 40 by wireless communication.

The CPU 113, the ROM 114, the RAM 115, the HDD 116, the display driver117, the operation input detector 118, and the wireless communicationunit 119 are connected to each other via a system bus BUS. Consequently,the CPU 113 can access the ROM 114, the RAM 115, and the HDD 116, cancontrol the display of various types of information on the display 111via the display driver 117, and can control the transmission andreception of various types of information to and from the radiationgenerator 120 and the electronic cassette 40 via the wirelesscommunication unit 119. Further, the CPU 113 can grasp states ofoperation by a user with respect to the operation panel 112 via theoperation input detector 118.

The radiation generator 120 includes the radiation source 121, awireless communication unit 123 that transmits and receives varioustypes of information such as the exposure conditions to and from theconsole 110, and a radiation source controller 122 that controls theradiation source 121 on the basis of the received exposure conditions.

The radiation source controller 122 is also configured to include amicrocomputer and stores the received exposure conditions and so forth.The exposure conditions received from the console 110 includeinformation such as tube voltage, tube current, and so forth. Theradiation source controller 122 causes the radiation X to be appliedfrom the radiation source 121 on the basis of the received exposureconditions.

Next, the action of the imaging system 104 pertaining to the presentembodiment will be described.

First, the action of the console 110 when capturing a radiographic imagewill be described with reference to FIG. 11. FIG. 11 is a flowchartshowing a flow of processing by a radiographic imaging processingprogram executed by the CPU 113 of the console 110 when an instructionto execute the program has been input via the operation panel 112; thisprogram is stored beforehand in a predetermined region of the ROM 114.

In step 300 of FIG. 11, the CPU 113 controls the display driver 117 soas to cause the display 111 to display a predetermined initialinformation input screen, and in the next step 302, the CPU 113 waitsfor the input of predetermined information.

In FIG. 12, there is shown an example of the initial information inputscreen displayed by the display 111 by the processing of step 300. Asshown in FIG. 12, in the initial information input screen pertaining tothe present embodiment, a message prompting the user to input the nameof the subject on which radiographic imaging is to be performed, thepart to be imaged, the posture during imaging, and the exposureconditions for exposure to the radiation X during imaging (in thepresent embodiment, the tube voltage and the tube current when exposingthe subject to the radiation X), and input fields for inputting thesepieces of information, are displayed.

When the initial information input screen shown in FIG. 12 is displayedon the display 111, the radiographer inputs the name of the subject tobe imaged, the part to be imaged, the posture during imaging, and theexposure conditions into the corresponding input fields via theoperation panel 112.

Then, in a case where the posture during imaging is the upright positionor the recumbent position, the radiographer puts the electronic cassette40 into the corresponding holder 162 of the upright-position stand 160or holder 166 of the recumbent-position table 164, positions theradiation source 121 in the corresponding position, and thereafterpositions the subject in a predetermined imaging position. With respectto this, in a case where the part to be imaged is an arm, a leg, or thelike and imaging is to be performed without putting the electroniccassette 40 into a holder, the radiographer positions the subject, theelectronic cassette 40, and the radiation source 121 so that the part tobe imaged is capable of being imaged. Thereafter, the radiographerdesignates, via the operation panel 112, the end button displayed in theneighborhood of the lower end of the initial information input screen.When the end button is designated by the radiographer, the determinationin step 302 becomes YES and the processing moves to step 304.

In step 304, the information (hereinafter called “initial information”)that has been input on the initial information input screen istransmitted to the electronic cassette 40 via the wireless communicationunit 119, and thereafter, in the next step 306, the exposure conditionsincluded in the initial information is transmitted to the radiationgenerator 120 via the wireless communication unit 119 to thereby set theexposure conditions. In response to this, the radiation sourcecontroller 122 of the radiation generator 120 prepares for exposure inthe received exposure conditions.

In the next step 308, the instruction information instructing the startof exposure is transmitted to the radiation generator 120 and theelectronic cassette 40 via the wireless communication unit 119.

In response to this, the radiation source 121 starts emitting theradiation X at the tube voltage and tube current corresponding to theexposure conditions that the radiation generator 120 received from theconsole 110. The radiation X emitted from the radiation source 121passes through the subject and thereafter reaches the electroniccassette 40.

When the cassette controller 58 of the electronic cassette 40 receivesthe instruction information instructing the start of exposure, thecassette controller 58 acquires the radiation dose information by meansof the radiation dose acquisition function and stands by until theradiation dose indicated by the acquired radiation dose informationbecomes equal to or greater than a first threshold value predeterminedas a value for detecting that the application of the radiation has beenstarted. Next, after the electronic cassette 40 has started the actionof capturing the radiographic image, the electronic cassette 40 stopsthe imaging action, and transmits exposure stop information to theconsole 110, at the point in time when the cumulative value of theradiation dose indicated by the radiation dose information has reached asecond threshold value predetermined as a value for stopping theexposure to the radiation X on the basis of the part to be imaged andthe exposure conditions included in the initial information.

Thus, in the next step 310, the CPU 113 waits to receive the exposurestop information, and in the next step 312, the instruction informationinstructing the stopping of the exposure to the radiation X istransmitted to the radiation generator 120 via the wirelesscommunication unit 119. In response to this, the exposure to theradiation X from the radiation source 121 is stopped.

When the electronic cassette 40 stops the action of capturing theradiographic image, the electronic cassette 40 transmits to the console110 the image data obtained by the imaging.

Thus, in the next step 314, the CPU 113 stands by until the image dataare received from the electronic cassette 40, and in the next step 316,the defective pixel correction processing is performed with respect tothe received image data and thereafter image processing that performsvarious types of correction such as shading correction is executed.

In the next step 318, the image data on which the image processing hasbeen performed (hereinafter called “corrected image data”) is stored inthe HDD 116, and in the next step 320, the display driver 117 iscontrolled so as to cause the display 111 to display the radiographicimage represented by the corrected image data for checking and so forth.

In the next step 322, the corrected image data is transmitted to the RISserver 150 via the in-hospital network 102, and thereafter theradiographic imaging processing program ends. The corrected image datathat have been transmitted to the RIS server 150 are stored in thedatabase 150A so that it becomes possible for doctors to read thecaptured radiographic image and make a diagnosis.

Next, the action of the electronic cassette 40 upon having received theinitial information from the console 110 will be described withreference to FIG. 13. FIG. 13 is a flowchart showing a flow ofprocessing by a cassette imaging processing program executed by the CPU58A in the cassette controller 58 of the electronic cassette 40 at thistime; this program is stored beforehand in a predetermined region of thememory 58B.

In step 400 of FIG. 13, the CPU 58A waits to receive the instructioninformation instructing the start of exposure from the console 110, andin the next step 402, a first radiation dose acquisition processingroutine program that acquires the radiation dose information by means ofthe radiation dose acquisition function is executed.

The first radiation dose acquisition processing routine programpertaining to the present embodiment will be described below withreference to FIG. 14. FIG. 14 is a flowchart showing a flow ofprocessing by the first radiation dose acquisition processing routineprogram pertaining to the present embodiment; this program is alsostored beforehand in a predetermined region of the memory 58B.

In step 500 of FIG. 14, the electric charges that had been stored in thecapacitors 92B is discharged by switching the reset switches 92F of allthe variable gain pre-amps 92 and 92′ to an on-state for a predeterminedperiod of time and the second signal processing unit 55 is reset bydischarging the signal levels held in all the sample-and-hold circuits97.

In the next step 502, all the sample-and-hold circuits 97 correspondingto the radiation detection pixels 32A of the peripheral edge portiondetection regions 20C to 20F (hereinafter called “peripheral edgeportion pixels”) are driven for a predetermined period of time tothereby cause the sample-and-hold circuits 97 to hold the signal levelsof the filtered electrical signals. In the next step 504, themultiplexer 98 is controlled in such a way that the output signals fromthe sample-and-hold circuits 97 corresponding to the peripheral edgeportion pixels are sequentially selected and output.

Because of the above processing, after being amplified by the variablegain pre-amps 92 or the variable gain pre-amps 92′, the electricalsignals corresponding to the peripheral edge portion pixels of theperipheral edge portion detection regions 20C and 20F are synthesized bythe binning components 94, and digital data representing the signallevels of the electrical signals filtered by the LPFs 96 aresequentially output from the A/D converter 99 as the radiation doseinformation. In the next step 506, the radiation dose information outputfrom the A/D converter 99 is sequentially acquired and thereafter thefirst radiation dose acquisition processing routine program ends. Whenthe first radiation dose acquisition processing routine program ends,the processing moves to step 404 of the cassette imaging processingprogram (main routine) shown in FIG. 13.

In step 404, it is determined whether or not the radiation dose (in thepresent embodiment, the average value of the radiation dose representedby the radiation dose information sequentially output from the A/Dconverter 99) represented by the information acquired by the processingof step 402 is equal to or greater than the first threshold value. In acase where the determination is NO in step 404, the processing returnsto step 402, and in a case where the determination is YES, this isregarded to mean that exposure to the radiation X from the radiationsource 121 has been started and moves to step 406.

In step 406, the electric charges stored in the capacitors 9 in thepixels 32 of the radiation detector 20 is discharged and thereafter thestorage of the electric charges in the capacitors 9 starts again tothereby start the action of capturing the radiographic image. In thenext step 408, a second radiation dose acquisition processing routineprogram that acquires the radiation dose information is executed by theradiation dose acquisition function.

The second radiation dose acquisition processing routine programpertaining to the present embodiment will be described below withreference to FIG. 15. FIG. 15 is a flowchart showing a flow ofprocessing by the second radiation dose acquisition processing routineprogram pertaining to the present embodiment; this program is alsostored beforehand in a predetermined region of the memory 58B.

In step 550 of FIG. 15, the electric charges that had been stored in thecapacitors 92B is discharged by switching the reset switches 92F of allthe variable gain pre-amps 92 and 92′ to an on-state for a predeterminedperiod of time and the second signal processing unit 55 is reset bydischarging the signal levels held in all the sample-and-hold circuits97.

In the next step 552, all the sample-and-hold circuits 97 correspondingto the radiation detection pixels 32A of the central portion detectionregions 20A and 20B (hereinafter called “center portion pixels”) aredriven for a predetermined period of time to thereby cause thesample-and-hold circuits 97 to hold the signal levels of the filteredelectrical signals. In the next step 554, the multiplexer 98 iscontrolled in such a way that the output signals from thesample-and-hold circuits 97 corresponding to the center portion pixelsare sequentially selected and output.

Because of the above processing, digital data representing the signallevels of the electrical signals filtered by the LPFs 96 or LPFs 96′after being amplified by the variable gain pre-amps 92 are sequentiallyoutput from the A/D converter 99 as the radiation dose information. Inthe next step 556, the radiation dose information output from the A/Dconverter 99 is sequentially acquired and thereafter the secondradiation dose acquisition processing routine program ends. When thesecond radiation dose acquisition processing routine program ends, theprocessing moves to step 410 of the cassette imaging processing program(main routine) shown in FIG. 13.

In step 410, it is determined whether or not the radiation dose (in thepresent embodiment, the average value of the radiation dose representedby the radiation dose information sequentially output from the A/Dconverter 99) represented by the information acquired by the processingof step 408 is equal to or greater than the second threshold value. In acase where the determination is NO in step 410, the processing moves tostep 412, the radiation dose acquired by the processing of step 408 iscumulated, the processing returns to step 408, and moves to step 414 atthe point in time when the determination becomes YES. When repeatedlyexecuting the processing of step 408 to step 412, in step 410 it isdetermined whether or not the radiation dose cumulated up until then hasbecome equal to or greater than the second threshold value.

In step 414, the imaging action started by the processing of step 406stops. In the next step 416, the exposure stop information istransmitted to the console 110 via the wireless communication unit 60.

In the next step 418, the gate line driver 52 is controlled so as tocause ON signals to be output sequentially one line at a time from thegate line driver 52 to the gate lines 34, and the thin-film transistors10 connected to the gate lines 34 are sequentially switched on one lineat a time.

In the radiation detector 20, when the thin-film transistors 10connected to the gate lines 34 are sequentially switched on one line ata time, the electric charges stored in the capacitors 9 flow out to thedata lines 36 as electrical signals sequentially one line at a time. Theelectrical signals that have flowed out to the data lines 36 areconverted into digital image data by the first signal processing unit54, and the digital image data are stored in the image memory 56.

Thus, in step 418, the image data stored in the image memory 56 is readout, and in the next step 420, the read image data is transmitted to theconsole 110 via the wireless communication unit 60 and thereafter thecassette imaging processing program ends.

Incidentally, as shown in FIG. 8, the radiation detector 20 is builtinto the electronic cassette 40 pertaining to the present embodiment insuch a way that the radiation X is applied from the TFT substrate 30side.

Here, as shown in FIG. 16, in a case where the radiation detector 20 isconfigured as a so-called penetration side sampling type in which theradiation is applied from the side on which the scintillator 8 is formedand the radiographic image is read by the TFT substrate 30 disposed onthe back surface side of the surface on which the radiation is madeincident, light is emitted more strongly on the upper surface side ofthe scintillator 8 in FIG. 16 (the opposite side of the TFT substrate 30side). Further, in a case where the radiation detector 20 is configuredas a so-called irradiation side sampling type in which the radiation isapplied from the TFT substrate 30 side and the radiographic image isread by the TFT substrate 30 disposed on the front surface side of thesurface on which the radiation is made incident, the radiation that haspassed through the TFT substrate 30 is made incident on the scintillator8, and the TFT substrate 30 side of the scintillator 8 more stronglyemits light. In the sensor portions 13 disposed on the TFT substrate 30,electric charges are generated by the light generated by thescintillator 8. For this reason, the emission position of thescintillator 8 with respect to the TFT substrate 30 is closer in a casewhere the radiation detector 20 is configured as an irradiation sidesampling type than in a case where the radiation detector 20 isconfigured as a penetration side sampling type, so the resolution of theradiographic image obtained by imaging is higher.

Further, in the radiation detector 20, the photoelectric conversion film4 is configured by an organic photoelectric conversion material, andvirtually no radiation is absorbed by the photoelectric conversion film4. For this reason, in the radiation detector 20 pertaining to thepresent embodiment, the amount of radiation absorbed by thephotoelectric conversion film 4 is small even in a case where theradiation passes through the TFT substrate 30 because of the irradiationside sampling type, so a drop in sensitivity with respect to theradiation can be suppressed. In the irradiation side sampling type, theradiation passes through the TFT substrate 30 and reaches thescintillator 8, but in a case where the photoelectric conversion film 4of the TFT substrate 30 is configured by an organic photoelectricconversion material in this way, there is virtually no absorption of theradiation by the photoelectric conversion film 4 and attenuation of theradiation can be kept small, so configuring the photoelectric conversionfilm 4 with an organic photoelectric conversion material is suited tothe irradiation side sampling type.

Further, the amorphous oxide configuring the active layers 17 of thethin-film transistors 10 and the organic photoelectric conversionmaterial configuring the photoelectric conversion film 4 are bothcapable of being formed into films at a low temperature. For thisreason, the substrate 1 can be formed by plastic resin, aramid, orbionanofibers in which there is little absorption of radiation. In thesubstrate 1 formed in this way, the amount of radiation absorbed issmall, so a drop in sensitivity with respect to the radiation can besuppressed even in a case where the radiation passes through the TFTsubstrate 30 because of the irradiation side sampling type.

Further, according to the present embodiment, as shown in FIG. 8, theradiation detector 20 is adhered to the top panel 41B inside the casing41 in such a way that the TFT substrate 30 is on the top panel 41B side,but in a case where the substrate 1 is formed by plastic resin, aramid,or bionanofibers whose rigidity is high, the rigidity of the radiationdetector 20 itself is high, so the top panel 41B of the casing 41 can beformed thin. Further, in a case where the substrate 1 is formed byplastic resin, aramid, or bionanofibers whose rigidity is high, theradiation detector 20 itself has flexibility, so the radiation detector20 does not easily break even in a case where shock has been applied tothe imaging region 41A.

As described in detail above, in the present embodiment, pluralradiation detection pixels (in the present embodiment, the radiationdetection pixels 32A) with mutually different characteristics arearranged in a radiation detector (in the present embodiment, theradiation detector 20), and the radiation detection pixels are usedselectively according to the characteristics to detect a state ofapplication of radiation, so the state of application of the radiationcan be detected more accurately compared to a case where the radiationdetector does not have these pixels.

Further, in the present embodiment, the radiation detection pixels arearranged in different positions in the imaging region, so the state ofapplication of the radiation can be detected more accurately as a resultof being able to select and use, in accordance with the size and soforth of the part to be imaged, the pixels used for detecting the stateof application of the radiation.

Further, in the present embodiment, the state of application of theradiation is the start of application of the radiation and the applieddose of the radiation, and the radiation detection pixels used in thedetection of the state of application are selected in accordance withthe state of application, so the state of application of the radiationcan be detected more accurately as a result of being able to select anduse, in accordance with the state of application, the pixels used todetect the state of application of the radiation.

Further, in the present embodiment, the characteristics are differentbecause of the plural radiation detection pixels being connected toamplifiers (in the present embodiment, the variable gain pre-amps 92 and92′) that amplify, at mutually different gains, signals represented bythe electric charges stored by the radiation detection pixels. Further,the characteristics are different because of the plural radiationdetection pixels being connected to low-pass filters (in the presentembodiment, the LPFs 96 and 96′) that perform low-pass filtering atmutually different low-pass filtering frequencies with respect tosignals represented by the electric charges stored by the radiationdetection pixels. Moreover, the characteristics are different because ofthe plural radiation detection pixels being connected to synthesizingmeans (in the present embodiment, the binning components 94) thatsynthesize mutually different numbers of signals represented by theelectric charges stored by the radiation detection pixels. Consequently,the characteristics can be realized easily.

Moreover, in the present embodiment, the radiation detector includesdedicated lines (in the present embodiment, the direct read-out lines38) for reading out the stored electric charges from the radiationdetection pixels, so the radiographic image can be captured at a higherspeed as a result of being able to detect the state of application ofthe radiation can be detected independently of the action of capturingthe radiographic image.

The present invention has been described above using an embodiment, butthe technical scope of the present invention is not limited to the scopedescribed in the embodiment. Various changes or improvements can be madeto the embodiment without departing from the gist of the invention, andthe technical scope of the present invention also includes embodimentsto which such changes or improvements have been made.

Further, the embodiment is not intended to limit the inventionspertaining to the claims, and it is not the case that all combinationsof features described in the embodiment are essential to the solution ofthe invention. The embodiment includes inventions of various stages, andvarious inventions can be extracted by appropriately combining theplural configural requirements disclosed. Even when several configuralrequirements are omitted from all the configural requirements disclosedin the embodiment, configurations from which those several configuralrequirements have been omitted can also be extracted as inventions aslong as effects are obtained.

For example, in the above embodiment, a case was described where, asshown in FIG. 6, the radiation detection pixels 32A are arrangedsymmetrically with respect to both the up-and-down direction and theleft-and-right direction in the central portion detection regions andthe peripheral edge portion detection regions, but the present inventionis not limited to this, and there are no particular limitations on thearranged positions of the radiation detection pixels 32A. However, byarranging the radiation detection pixels in such a way as to besymmetrical with respect to the up-and-down direction and theleft-and-right direction like in the present embodiment, the electroniccassette 40 can be used without having to worry about the up-and-downdirection of the electronic cassette 40, so user-friendliness can beimproved, which is preferred.

Here, in a case where the radiation detection pixels 32A are arranged insuch a way as to not be symmetrical with respect to the up-and-downdirection, it is preferred that the electronic cassette 40 be given aconfiguration in which direction detecting means such as an accelerationsensor or a gyro is disposed in the electronic cassette 40 and in whichthe positions of the radiation detection pixels 32A are identified inaccordance with the direction of the electronic cassette 40 identifiedby the direction detecting means.

In the above embodiment, some of the pixels 32 disposed in the radiationdetector 20 are used as the radiation detection pixels 32A, so needlessto say it is preferred that adjacent radiation detection pixels 32A bespaced apart from each other to an extent that the defective pixelcorrection can be implemented.

Further, in the above embodiment, the radiation detection pixels 32Awhose characteristics are different are dispersed and arranged by grouppositions where the characteristics in the imaging region of theradiation detector 20 can be optimally utilized, so the electroniccassette 40 may also be given a configuration where marks such ascharacters, symbols, or designs denoting the correspondingcharacteristics are disposed in positions on the front surface of thetop panel 41B of the electronic cassette 40 corresponding to thepositions in which the pixel groups of the radiation detection pixels32A are arranged, so that the radiographer may reference the marks andselect and use the radiation detection pixels 32A. In this case, theelectronic cassette 40 may be given a configuration in which the colorsof the marks are changed per characteristic.

Further, in the above embodiment, as shown in FIG. 6 and FIG. 10, a casewas described where the radiation detection pixels 32A are arranged insuch a way as to have different characteristics per arrangement region,but the present invention is not limited to this; for example, as shownin FIG. 17, the radiation detector 20 may also be given a configurationin which the radiation detection pixels 32A that have differentcharacteristics are arranged in the same arrangement regions.

Further, in the above embodiment, a case was described where some of thepixels 32A disposed in the radiation detector 20 are used as theradiation detection pixels 32A, but the present invention is not limitedto this; for example, the radiation detector 20 may also be given aconfiguration in which the radiation detection pixels 32A are layered inthe radiation detector 20 as a separate layer from the pixels 32. Inthis case, no defective pixels arise, so the quality of the radiographicimage can be improved compared to the above embodiment. Further, in thiscase, it becomes possible to ensure that the radiation detection pixelshave different characteristics depending on the light-receiving area,the constituent material, and so forth of the radiation detection pixels32A, and it becomes unnecessary to dispose the second signal processingunit 55 pertaining to the above embodiment.

Further, in the above embodiment, a case was described where theradiation detection pixels 32A are dedicated pixels that detect theradiation, but the present invention is not limited to this and may alsobe given a configuration that uses the radiation detection pixels 32Adoubly as the radiographic image acquisition pixels 32B.

Further, in the above embodiment, a case was described where just twotypes of gains of the variable gain pre-amps and two types of low-passfiltering frequencies of the LPFs were disposed, but the presentinvention is not limited to this and may also be given a configurationwhere three or more types of these are disposed. Further, the number ofthe electrical signals synthesized by the binning components is also notlimited to two, and the present invention may also be given aconfiguration where the number is three or more.

Further, in the above embodiment, a case was described where theradiation detection pixels 32A were used to detect the start ofapplication of the radiation and the applied dose, but the presentinvention is not limited to this and may also be given a configurationwhere the radiation detection pixels 32A are used to detect the stop ofapplication of the radiation.

Further, the applied conditions of the gains of the variable gainpre-amps, the binning state resulting from the binning components, andthe low-pass filtering frequencies of the LPFs described in the aboveembodiment are examples, and example configurations described below canalso be employed.

Regarding the gains of the variable gain pre-amps, a configuration thatapplies a lower gain the greater the applied dose of the radiation X is,a configuration that applies a higher gain in the case of capturing amoving image than in the case of capturing a still image, and aconfiguration that applies a relatively high gain in the case ofdetecting the start of the application of the radiation and applies arelatively low gain in the case of detecting the end of the applicationof the radiation and the applied dose can be exemplified.

Further, regarding the binning states resulting from the binningcomponents, a configuration that applies a lower binning number thegreater the applied dose of the radiation X is, a configuration thatapplies a higher binning number in the case of capturing a moving imagethan in the case of capturing a still image, and a configuration thatapplies a relatively high binning number in the case of detecting thestart of the application of the radiation and applies a relatively lowbinning number in the case of detecting the end of the application ofthe radiation and the applied dose can be exemplified.

Moreover, regarding the low-pass filtering frequencies of the LPFs, aconfiguration that applies a lower low-pass filtering frequency thelower the tube current and tube voltage when emitting the radiation Xare and a configuration that applies a relatively low low-pass filteringfrequency as the low-pass filtering frequency corresponding to theradiation detection pixels positioned in the imaging region where thepart to be imaged is positioned can be exemplified.

Further, in the above embodiment, a case was described where theradiation detection pixels 32A arranged side-by-side in the rowdirection in the radiation detector 20 are connected to common directread-out lines 38, but the present invention is not limited to this andmay also be given a configuration where all the radiation detectionpixels 32A are individually connected to different direct read-out lines38.

Further, in the above embodiment, a case was described where the sensorportions 13 are configured to include the organic photoelectricconversion material in which electric charge is generated as a result ofreceiving the light generated by the scintillator 8, but the presentinvention is not limited to this and may also be given a configurationthat applies sensor portions configured to not include the organicphotoelectric conversion material as the sensor portions 13.

Further, in the above embodiment, a case was described where theradiation detector 20 and the case 42 accommodating the cassettecontroller 58 and the power source 70 are placed inside the casing 41 ofthe electronic cassette 40 in such a way as to not coincide, but thepresent invention is not limited to this. For example, the cassettecontroller 58 and the power source 70 may also be placed in such a wayas to coincide with the radiation detector 20.

Further, in the above embodiment, a case was described wherecommunication is performed wirelessly between the electronic cassette 40and the console 110 and between the radiation generator 120 and theconsole 110, but the present invention is not limited to this and mayalso be given a configuration where, for example, communication betweenat least one of these is performed via wires.

Further, in the above embodiment, a case was described where X-rays areapplied as the radiation, but the present invention is not limited tothis and may also be given a configuration where another form ofradiation such as gamma rays is applied.

In addition, the configuration of the RIS 100 (see FIG. 1), theconfiguration of the radiographic imaging room (see FIG. 2), theconfiguration of the electronic cassette 40 (see FIG. 3 to FIG. 8 andFIG. 10), and the configuration of the imaging system 104 (see FIG. 9)described in the above embodiment are examples, and needless to sayunnecessary portions can be deleted therefrom, new portions can be addedthereto, and states of connection and so forth can be changed withoutdeparting from the gist of the present invention.

Further, the configuration of the initial information described in theabove embodiment is also an example, and needless to say unnecessaryinformation can be deleted therefrom and new information can be addedthereto without departing from the gist of the present invention.

Further, the flows of processing by the various programs (see FIG. 11and FIG. 13 to FIG. 15) described in the above embodiment are alsoexamples, and unnecessary steps can be deleted therefrom, new steps canbe added thereto, and the processing order can be switched aroundwithout departing from the gist of the present invention.

Further, the configuration of the initial information input screen (seeFIG. 12) described in the above embodiment is also an example, andunnecessary information can be deleted therefrom and new information canbe added thereto without departing from the gist of the presentinvention.

1. A radiographic imaging device comprising: a radiation detectorcomprising a plurality of pixels that are arranged in a matrix in animaging region that comprises a plurality of radiographic imageacquisition pixels for capturing a radiographic image and that acquireimage information representing the radiographic image by convertingapplied radiation into electric charges and storing the electriccharges, and a plurality of radiation detection pixels that havemutually different characteristics, and that detect the appliedradiation by converting the applied radiation into electric charges andstoring the electric charges; and a detecting unit that uses theradiation detection pixels selectively according to the mutuallydifferent characteristics to detect a state of application of theradiation.
 2. The radiographic imaging device according to claim 1,further comprising an acquiring unit that acquires an imaging conditionfor capturing the radiographic image, wherein the detecting unit usesthe radiation detection pixels which have characteristics correspondingto the imaging condition acquired by the acquiring unit to detect thestate of application of the radiation.
 3. The radiographic imagingdevice according to claim 1, wherein the radiation detection pixels arearranged in different positions in the imaging region.
 4. Theradiographic imaging device according to claim 1, wherein the state ofapplication of the radiation is at least one of a start of applicationof the radiation, an end of application of the radiation, or an theapplied dose of the radiation, and the detecting unit selects theradiation detection pixels used in the detection of the state ofapplication in accordance with the state of application of the radiationto be detected.
 5. The radiographic imaging device according to claim 1,wherein the mutually different characteristics are different because ofthe plurality of radiation detection pixels being connected toamplifiers that amplify, at mutually different gains, signalsrepresented by the electric charges stored by the radiation detectionpixels.
 6. The radiographic imaging device according to claim 1, whereinthe mutually different characteristics are different because of theplurality of radiation detection pixels being connected to low-passfilters that perform low-pass filtering at mutually different low-passfiltering frequencies with respect to signals represented by theelectric charges stored by the radiation detection pixels.
 7. Theradiographic imaging device according to claim 1, wherein the mutuallydifferent characteristics are different because of the plurality ofradiation detection pixels being connected to a synthesizing unit thatsynthesize mutually different numbers of signals represented by theelectric charges stored by the radiation detection pixels.
 8. Theradiographic imaging device according to claim 2, wherein the imagingcondition is at least one of a part to be imaged, a region in which thepart to be imaged is placed when capturing the radiographic image,whether the imaging is imaging to capture a moving image or a stillimage, or an applied dose of the radiation.
 9. The radiographic imagingdevice according to claim 1, wherein the radiation detector furthercomprises dedicated lines for reading out the stored electric chargesfrom the radiation detection pixels.
 10. A non-transitory computerreadable medium storing a program executed by a radiographic imagingdevice comprising a radiation detector comprising a plurality of pixelsthat are arranged in a matrix in an imaging region which comprises aplurality of radiographic image acquisition pixels for capturing aradiographic image and that acquire image information representing theradiographic image by converting applied radiation into electric chargesand storing the electric charges and a plurality of radiation detectionpixels that have mutually different characteristics, and that detect theapplied radiation by converting the applied radiation into electriccharges and storing the electric charges, the program causing a computerto function as an acquiring unit that acquires an imaging condition forcapturing the radiographic image and a detecting unit that uses theradiation detection pixels having characteristics corresponding to theimaging condition acquired by the acquiring unit to detect a state ofapplication of the radiation.
 11. The radiographic imaging deviceaccording to claim 5, wherein the mutually different characteristics aredifferent because of the plurality of radiation detection pixels beingconnected to low-pass filters that perform low-pass filtering atmutually different low-pass filtering frequencies with respect tosignals represented by the electric charges stored by the radiationdetection pixels.
 12. The radiographic imaging device according to claim5, wherein the mutually different characteristics are different becauseof the plurality of radiation detection pixels being connected to asynthesizing unit that synthesize mutually different numbers of signalsrepresented by the electric charges stored by the radiation detectionpixels.
 13. The radiographic imaging device according to claim 1,wherein the mutually different characteristics are different because ofthe plurality of radiation detection pixels being connected toamplifiers that amplify, at mutually different gains, signalsrepresented by the electric charges stored by the radiation detectionpixels, wherein the mutually different characteristics are differentbecause of the plurality of radiation detection pixels being connectedto low-pass filters that perform low-pass filtering at mutuallydifferent low-pass filtering frequencies with respect to signalsrepresented by the electric charges stored by the radiation detectionpixels, and wherein the mutually different characteristics are differentbecause of the plurality of radiation detection pixels being connectedto a synthesizing unit that synthesize mutually different numbers ofsignals represented by the electric charges stored by the radiationdetection pixels.